Image-forming MR methods which utilize the interaction between magnetic fields and nuclear spins in order to form two-dimensional or three-dimensional images are widely used nowadays, notably in the field of medical diagnostics, because for the imaging of soft tissue they are superior to other imaging methods in many respects, do not require ionizing radiation and are usually not invasive.
According to the MR method in general, the body of the patient to be examined is arranged in a strong, uniform magnetic field (B0 field) whose direction at the same time defines an axis (normally the z-axis) of the co-ordinate system on which the measurement is based. The magnetic field produces different energy levels for the individual nuclear spins in dependence on the magnetic field strength which can be excited (spin resonance) by application of an electromagnetic alternating field (RF field, also referred to as B1 field) of defined frequency (so-called Larmor frequency, or MR frequency). From a macroscopic point of view the distribution of the individual nuclear spins produces an overall magnetization which can be deflected out of the state of equilibrium by application of an electromagnetic pulse of appropriate frequency (RF pulse) while the magnetic field extends perpendicular to the z-axis, so that the magnetization performs a precessional motion about the z-axis. The precessional motion describes a surface of a cone whose angle of aperture is referred to as flip angle. The magnitude of the flip angle is dependent on the strength and the duration of the applied electromagnetic pulse. In the case of a so-called 90° pulse, the spins are deflected from the z axis to the transverse plane (flip angle 90°).
After termination of the RF pulse, the magnetization relaxes back to the original state of equilibrium, in which the magnetization in the z direction is built up again with a first time constant T1 (spin lattice or longitudinal relaxation time), and the magnetization in the direction perpendicular to the z direction relaxes with a second time constant T2 (spin-spin or transverse relaxation time). The variation of the magnetization can be detected by means of one or more receiving RF coils which are arranged and oriented within an examination volume of the MR device in such a manner that the variation of the magnetization is measured in the direction perpendicular to the z-axis. The decay of the transverse magnetization is accompanied, after application of, for example, a 90° pulse, by a transition of the nuclear spins (induced by local magnetic field inhomogeneities) from an ordered state with the same phase to a state in which all phase angles are uniformly distributed (dephasing). The dephasing can be compensated by means of a refocusing pulse (for example a 180° pulse). This produces an echo signal (spin echo) in the receiving coils.
In order to realize spatial resolution in the body, linear magnetic field gradients extending along the three main axes are superposed on the uniform magnetic field, leading to a linear spatial dependency of the spin resonance frequency. The signal picked up in the receiving coils then contains components of different frequencies which can be associated with different locations in the body. The MR signal data obtained via the RF coils corresponds to the spatial frequency domain and is called k-space data. The k-space data usually includes multiple lines acquired with different phase encoding. Each line is digitized by collecting a number of samples. A set of k-space data is converted to a MR image by means of Fourier transformation.
In a number of MR-guided diagnostic and therapeutic procedures, the measurement of tissue temperature is of particular importance. Thus, for example, in MR-HIFU (‘magnetic resonance-guided high intensity focused ultrasound’) applications, in which tissue is locally heated and destroyed by the focused irradiation of ultrasound energy, the local change of the temperature distribution needs to monitored during the procedure in order to be able to control the irradiation of ultrasound in a targeted manner. The heating regime must be adapted to keep the local temperature increase for the healthy tissue within allowable margins, while the target to be destroyed (for example a malignant tumor) is sufficiently heated. Besides MR-HIFU there are a number of other therapeutic procedures as well as diagnostic and functional MR studies, in which the spatial distribution of a change in temperature may be of interest and needs to be monitored.
One of the most sensitive MR-based temperature mapping approaches is the known proton resonance frequency shift (PRF) method (Rieke et al, Magnetic Resonance in Medicine, volume 51, pages 1223-1231). The magnetic resonance frequency of water protons changes as a function of temperature. Temperature changes induce slight variations of the bonding angles between the protons in the water molecules resulting in variations of the electronic shielding, resulting in a small change of the chemical shift. For a gradient echo acquisition performed at a given echo-time at two different temperatures a local change in the signal phase can be observed. A drawback is that the PRF method requires rather long echo times (in the order of T2*) for optimal sensitivity. Therefore, to allow real-time temperature mapping, typically spiral- or EPI-based fast MR signal readouts are used, which are prone to different kinds image artifacts (ghosting, blurring, image distortions, chemical shift-related artifacts etc.). Moreover, temperature maps acquired by the conventional PRF method, are prone to errors induced by different kinds of effects influencing the MR signal phase, such as patient/organ motion, overall system drift (RF, main field magnetic field, heating of the gradient coils etc.), eddy currents and so forth. Furthermore, the PRF method can be compromised by the signal composition within each individual voxel of the acquired MR image. Fat does not show the PRF effect. This can result in erroneous temperature change estimates in case both fat protons and water protons contribute to the MR signal within a single voxel. This is a particular problem in the context of MR-HIFU because MR-HIFU is typically applied in the abdominal chamber, where significant amounts of visceral fat may be present. Finally, flow represents a problem in PRF methods because flow-related phase artifacts may degrade the accuracy of the PRF shift measurements.
The European patent application EP 2 615 470 discloses that a stimulated echo acquisition preparation sequence is employed to encode phase-information. This approach is employed to encode the local flip angle from which the B1-field distribution is derived.